System for electrosurgical myocardial revascularization

ABSTRACT

A method for transmyocardial revascularization of the heart of a patient includes positioning an active electrode surface in close proximity to a target site on the wall of a patient&#39;s heart, and applying high frequency voltage between the active voltage surface and a return electrode to ablate tissue at the heart wall. The high frequency voltage ablates, i.e. volumetrically removes the heart tissue, and the electrode surface is axially translated into the space vacated by the removed tissue to bore a channel through the heart tissue. The active electrode surface may be introduced into the thoracic cavity and placed adjacent the epicardium to form an inward channel toward the ventricular cavity, or it may be delivered into the ventricular cavity of the heart and positioned adjacent the endocardium to form a channel extending outward towards the epicardium. In either case, the channels formed through the myocardium promote direct communication between blood within the ventricular cavity and that of existing myocardial vasculature to increase blood flow to the heart tissue.

The present invention is a divisional of application Ser. No. 09/054,660filed Apr. 3, 1998 now abandoned, which is a continuation of applicationSer. No. 08/753,227 filed Nov. 22, 1996, now U.S. Pat. No. 5,873,855,which is a continuation-in-part of application Ser. No. 08/562,331 filedNov. 22, 1995, now U.S. Pat. No. 5,683,366 which is acontinuation-in-part of application Ser. No. 08/485,219 filed Jun. 7,1995, now U.S. Pat. No. 5,697,281, which was a continuation-in-part ofPCT International Application, US National Phase Serial No.PCT/US94/05168 filed May 10, 1994, which was a continuation-in-part ofU.S. patent application Ser. No. 08/059,681, filed on May 10, 1993, nowabandoned, which was a continuation-in-part of U.S. application Ser. No.07/958,977, filed on Oct. 9, 1992, now U.S. Pat. No. 5,366,443 which isa continuation-in-part of U.S. application Ser. No. 07/817,575, filed onJan. 7, 1992, now abandoned, the full disclosures of which areincorporated herein by reference for all purposes.

BACKGROUND OF THE INVENTION

Field of the Invention

The present invention relates generally to the field of electrosurgeryand, more particularly, to surgical devices and methods that employ highfrequency energy to cut and ablate heart tissue for increasing the flowof blood to a patient's heart.

Coronary artery disease, the build up of atherosclerotic plaque on theinner walls of the coronary arteries, causes the narrowing or completeclosure of these arteries resulting in insufficient blood flow to theheart. A number of approaches have been developed for treating coronaryartery disease. In less severe cases, it is often sufficient to treatthe symptoms with pharmaceuticals and lifestyle modification to lessenthe underlying causes of the disease. In more severe cases a coronaryartery blockage can often be treated using endovascular techniques, suchas balloon angioplasty, a laser recanalization, placement of stents, andthe like.

In cases where pharmaceutical treatment and endovascular approaches havefailed or are likely to fail, it is often necessary to perform acoronary artery bypass graft procedure using open or thoracoscopicsurgical methods. For example, many patients still require bypasssurgery due to such conditions as the presence of extremely diffusestenotic lesions, the presence of total occlusions and the presence ofstenotic lesions in extremely tortuous vessels. However, some patientsare too sick to successfully undergo bypass surgery. For other patients,previous endovascular and/or bypass surgery attempts have failed toprovide adequate revascularization of the heart muscle.

The present invention is particularly concerned with an alternative tothe above procedures, which is known as laser myocardialrevascularization (LMR). LMR is a recent procedure developed with therecognition that myocardial circulation occurs through arterioluminalchannels and myocardial sinusoids in the heart wall, as well as throughthe coronary arteries. In LMR procedures, artificial channels are formedin the myocardium with laser energy to provide blood flow to ischemicheart muscles by utilizing the heart's ability to perfuse itself fromthese artificial channels through the arterioluminal channels andmyocardial sinusoids. In one such procedure, a CO₂ laser is utilized tovaporize tissue and produce channels in the heart wall from theepicardium through the endocardium to promote direct communicationbetween blood within the ventricular cavity and that of existingmyocardial vasculature. The laser energy is typically transmitted fromthe laser to the epicardium by an articulated arm device. Recently, apercutaneous method of LMR has been developed in which an elongatedflexible lasing apparatus is attached to a catheter and guidedendoluminally into the patient's heart. The inner wall of the heart isirradiated with laser energy to form a channel from the endocardium intothe myocardium for a desired distance.

While recent techniques in LMR have been promising, they also sufferfrom a number of drawbacks inherent with laser technology. One suchdrawback is that the laser energy must be sufficiently concentrated toform channels through the heart tissue, which reduces the diameter ofthe channels formed by LMR. In addition, free beam lasers generally mustcompletely form each artificial lumen or revascularizing channel duringthe still or quiescent period of the heart beat. Otherwise, the laserbeam will damage surrounding portions of the heart as the heart beatsand thus moves relative to the laser beam. Consequently, the surgeonmust typically form the channel in less than about 0.08 seconds, whichrequires a relatively large amount of energy. This further reduces thesize of the channels that may be formed with a given amount of laserenergy. Applicant has found that the diameter or minimum lateraldimension of these artificial channels may have an effect on theirability to remain open. Thus, the relatively small diameter channelsformed by existing LMR procedures (typically on the order of about 1 mmor less) may begin to close after a brief period of time, which reducesthe blood flow to the heart tissue.

Another drawback with current LMR techniques is that it is difficult toprecisely control the location and depth of the channels formed bylasers. For example, the speed in which the revascularizing channels areformed often makes it difficult to determine when a given channel haspierced the opposite side of the heart wall. In addition, the distancein which the laser beam extends into the heart is difficult to control,which can lead to laser irradiation with heating or vaporization ofblood or heart tissue within the ventricular cavity. For example, whenusing the LMR technique in a pericardial approach (i.e., from outsidesurface of the heart to inside surface), the laser beam may not onlypierce through the entire wall of the heart but may also irradiate bloodwithin the heart cavity. As a result, one or more blood thromboses orclots may be formed which can lead to vascular blockages elsewhere inthe circulatory system. Alternatively, when using the LMR technique inan endocardial approach (i.e., from the inside surface of the hearttoward the outside surface), the laser beam may not only pierce theentire wall of the heart but may also irradiate and damage tissuesurrounding the outer boundary of the heart.

2. Description of the Background Art

Devices incorporating radio frequency electrodes for use inelectrosurgical and electrocautery techniques are described in Rand etal. (1985) J. Arthro. Surg. 1:242-246 and U.S. Pat. Nos. 5,281,216;4,943,290; 4,936,301; 4,593,691; 4,228,800; and 4,202,337. U.S. Pat.Nos. 4,943,290 and 4,036,301 describe methods for injectingnon-conducting liquid over the tip of a monopolar electrosurgicalelectrode to electrically isolate the electrode, while energized, from asurrounding electrically conducting irrigant. U.S. Pat. Nos. 5,195,959and 4,674,499 describe monopolar and bipolar electrosurgical devices,respectively, that include a conduit for irrigating the surgical site.

U.S. Pat. Nos. 5,380,316, 4,658,817, 5,389,096, PCT Application No. WO94/14383, European Patent Application No. 0 515 867, and Articles“Transmyocardio Laser Revascularization”, Mirhoseini et al., Journal ofClinical Laser Medicine & Surgery Vol. 11, No. 1:15-19 (1993); “NewConcepts in Revascularization of the Myocardium”, Mirhoseini, et al.,The Annuals of Thoracic Surgery Society of Thoracic Surgeons, Vol. 45,No. 4:415-420 (1988); “Transmyocardial Acupuncture”, Sen, et al. Journalof Thoracic and Cardiovascular Surgery, Vol. 50, No. 2:181-189 (1965);“Transmural Channels Can Protect Ischemic Tissue”, Whittaker, et al.Circulation, Vol. 93, No. 1:143-152 (1996); “Regional myocardial bloodflow and cardiac mechanics in dog hearts with CO₂ laser-inducedintramyocardial revascularization”, Hardy, et al., Basic Res. Cardiol,85:179-196 (1990); “Treatment of Acute Myocardial Infarction byTransmural Blood Supply From the Ventricular Cavity”, Walter, et al.,Europ. Sure. Res., 130-138 (1971); “Revascularization of the Heart byLaser”, Mirhoseini and Clayton, Journal of Microsurgery, 2:253-260(1981); “Transventricular Revascularization by Laser”, Mirhoseini, etal., Lasers in Surgery and Medicine, 2:187-198 (1982) describe methodsand apparatus for percutaneous myocardial revascularization. Thesemethods and apparatus involve directing laser energy against the hearttissue to form transverse channels through the myocardium to increaseblood flow from the ventricular cavity to the myocardium.

SUMMARY OF THE INVENTION

The present invention provides systems, apparatus and methods forselectively applying electrical energy to structures within or on thesurface of a patient's body. The present invention allows the surgicalteam to perform electrosurgical interventions, such as ablation andcutting of body structures, while limiting the depth of necrosis andlimiting damage to tissue adjacent the treatment site. The systems,apparatus and methods of the present invention are particularly usefulfor canalizing or boring channels or holes through tissue, such as theventricular wall of the heart during transmyocardial revascularizationprocedures.

In a method according to the present invention, an active electrodesurface is positioned in close proximity to a target site on the wall ofa patient's heart, and high frequency voltage is applied between theactive voltage surface and a return electrode to ablate tissue at theheart wall. The high frequency voltage ablates, i.e. volumetricallyremoves the heart tissue, and the electrode surface is axiallytranslated into the space vacated by the removed tissue to bore achannel through the heart tissue. The active electrode surface may beintroduced into the thoracic cavity and placed adjacent the outer heartwall or epicardium to form an inward channel toward the ventricularcavity, or it may be delivered into the ventricular cavity of the heartand positioned adjacent the inner heart wall or endocardium to form achannel extending outward towards the epicardium. In either case, thechannels formed through the heart wall promote direct communicationbetween blood within the ventricular cavity and that of existingmyocardial vasculature to increase blood flow to the heart tissue.

One of the advantages of the present invention, particularly overprevious methods involving lasers, is that the surgeon can moreprecisely control the location, depth and diameter of therevascularizing channels formed in the heart tissue. For example, theactive electrode surface remains in contact with the heart wall as thehigh frequency voltage ablates the heart tissue (or at leastsubstantially close to the heart wall, e.g., usually on the order ofabout 0.1 to 2.0 mm and preferably about 0.1 to 1.0 mm). This preservestactile sense and allows the surgeon to more accurately determine whento terminate cutting of a given channel so as to minimize damage tosurrounding tissues and/or minimize bleeding into the thoracic cavity.In addition, axially translating the active electrode through the heartwall allows the surgeon to form the channel at a slower pace thanconventional LMR because the channel does not have to be completelyformed during the quiescent or diastolic period of the heart. Since theactive electrode array generally directs tissue ablating energy onlyabout 0.1 to 3.0 mm in front of the electrode array (and preferably onlyabout 0.1 to 2.0 mm in front of the electrode array), this relativelyslow ablation pace allows the surgeon to more accurately control thechannel depth.

In one embodiment, an electrosurgical probe having one or moreelectrodes on its distal end is delivered into the thoracic cavityexterior to the heart wall. The probe may be delivered directly througha median thoracotomy or through an intercostal percutaneous penetration,such as a cannula or trocar sleeve in the chest wall between twoadjacent ribs. The electrode or electrode array is then positioned inclose proximity to the epicardium in the region of the heart to becanalized, and a high frequency voltage is applied between the electrodeor electrode array and a return electrode to form artificial channelsthrough the heart tissue. The return electrode may be integral with theprobe. By way of example, the return electrode may be located on theperimeter of the probe shaft proximal to the ablating (active) electrodeor electrode array. In another embodiment, two or more electrodes ofopposite polarity may be positioned at the distal end of theelectrosurgical probe to effect ablation of the wall of the heart.Alternatively, the return electrode may be positioned on anotherinstrument that is, for example, delivered through the same or anotherintercostal trocar sleeve. The probe is axially translated through theartificial channel provided by the trocar sleeve as the active electrodeablates tissue to maintain contact with the heart wall and to facilitateprecise control of the procedure by the surgeon.

In another embodiment, the electrode array is introduced through apercutaneous penetration in the patient and axially translated throughone of the major arterial vessels to the left ventricular cavity. Inthis embodiment, the electrode or electrode array may form a distalportion of an electrosurgical catheter and may be guided through aconventional or specialized guide catheter. The electrode array is thenpositioned adjacent the endocardium and axially translated outward toform one or more channels through the myocardium. The surgeon maycontrol the depth of the channels by axially translating the catheterthrough the heart wall, and terminating the electrical energy to theactive array when the channel has reached the desired depth. Thechannels may be formed completely through the myocardium to the outersurface of the epicardium, or the surgeon may terminate the electricalenergy prior to penetrating the outer surface of the epicardium toprevent blood from flowing into the thoracic cavity.

The control of the depth of channel formed in the wall of the heart maybe accomplished using one or a combination of several methods andapparatus. By way of example, real-time fluoroscopic visualization ofthe heart in combination with radiopaque markers on the electrosurgicalcatheter may be used by the surgeon to control the depth of the channeland terminate ablation before penetrating through the outer surface ofthe heart wall. Alternatively, ultrasound methods may be incorporatedwithin the electrosurgical catheter or guide tube to determine thethickness of the heart wall adjacent to the distal probe tip and allowthe surgeon to pre-set the depth of each channel before energizing theprobe and ablating the heart tissue. Also, ultrasound methods may beincorporated within the electrosurgical catheter to continuously detectthe distance of the electrode or electrode array from the outer surfaceof the heart and to interrupt the voltage applied to the ablatingelectrode(s) in order to stop the forward advance of the catheter at apredetermined distance from the outer surface of the heart. In yetanother embodiment, the electrosurgical catheter includes a smalldiameter tissue electrical impedance measurement sensor (e.g., 0.1 to0.5 mm diameter) which extends distal to the tissue ablating electrodeor electrode array (e.g., 1 to 10 mm). This impedance measurement sensordetects the outer surface of the heart as it penetrates through thetissue and enters a region of different electrical impedance (i.e., thefluid-filled cavity surrounding the heart).

In another aspect of the present invention, radially expandable luminalprotheses, such as stents and stent-grafts, are implanted in one or moreof the revascularizing channels after the channels have been formed bythe electrosurgical instrument. The stents may be implanted immediatelyafter the channels have been formed (i.e., with the electrosurgicalprobe or catheter), or they may be implanted after the channels havebeen formed with a separate delivery catheter. The stents are compressedinto a narrow-diameter configuration, and advanced endoluminally to thetarget site in the heart tissue with a delivery catheter. Theintraluminal prostheses will typically comprise a resilient, radiallycompressible, tubular frame having a proximal end, a distal end, and anaxial lumen therebetween. The tubular frame includes a plurality ofopenings or slots that allow it to be expanded radially outward withinthe channel by conventional methods, such as shape memory alloys,expandable balloons, and the like. The stent exerts a radial forceagainst the inner channel walls to maintain patency of the channels,thereby increasing the blood flow from the ventricular cavity to themyocardium. In the case of stent-grafts, a porous liner, typically afabric, polymeric sheet, membrane, or the like, will line all or most ofthe luminal surface of the tubular frame to inhibit occlusion of thechannel through the openings in the tubular frame while allowingoxygenated blood to pass through the porous liner and into the hearttissues surrounding the channel.

The apparatus according to the present invention comprises anelectrosurgical instrument having a shaft with a proximal end, a distalend and one or more active electrodes at or near the distal end. Areturn electrode is disposed on the shaft close to the distal end and aconnector extends through the shaft for electrically coupling the activereturn electrodes to a high frequency voltage source. The distal portionof the shaft and the active electrodes are sized for delivery through atrocar canalization (e.g., pericardial approach) or guiding catheter(e.g., endocardial approach) to ablate tissue in the heart wall to forma revascularizing channel through at least a portion of the heart wall.The return electrode may be provided integral with the shaft, or it maybe separate from the shaft.

The shaft may also incorporate means for delivery of electricallyconductive liquid (e.g., isotonic saline) to the distal end of theelectrosurgical instrument to provide an electrically conductive pathwaybetween the one or more active electrodes and the return electrode. Theelectrosurgical instrument may also include an ultrasonic transducer foreither measuring the thickness of the heart wall (for pre-setting thedepth of canalization) or detecting the distance from the distal end ofthe electrosurgical instrument to the outer surface of the heart tointerrupt the ablation of the heart wall (and depth of canalization) ata preselected distance from the outer surface of the heart wall usingactive feedback control within the power source. Alternatively, theelectrosurgical instrument may include an electrical impedance measuringsensor for detecting the distance form the distal end of theelectrosurgical instrument to the outer surface of the heart tointerrupt the ablation of the heart wall (and depth of canalization) ata preselected distance from the outer surface of the heart wall usingactive feedback control with the power source (e.g., when the measuredelectrical impedance at the tip of the sensor increases above apreselected level, the applied voltage is interrupted).

In an exemplary embodiment, the instrument comprises an electrosurgicalprobe having at least a distal end configured for delivery through anintercostal penetration in the patient, such as a trocar sleevepositioned between two ribs. The probe preferably includes an electrodearray with a plurality of isolated electrode terminals at its distalend. A return electrode is proximally recessed from the electrode arrayfor applying high frequency voltage therebetween to ablate or bore ahole through the heart tissue. The probe may include a fluid channel fordirecting electrically conducting fluid to the target site to completethe current return path from the heart tissue to the return electrode.Alternatively, this path may be completed by the heart tissue on theside of the probe, or the blood and other fluids existing within theheart wall.

In another embodiment, the electrosurgical instrument comprises a guidecatheter having a flexible, steerable shaft configured for endoluminaldelivery into the ventricular cavity. The guide catheter provides aninterior lumen through which an electrosurgical catheter can be deployedpercutaneously to form a channel in the wall of the heart. The guidecatheter is first positioned on the endocardial surface of the heart atthe site of a required channel. Next, the electrosurgical catheterlocated within the lumen of the guide catheter is positioned against thesurface of the endocardium and energized while advancing to apreselected channel depth based on one or a combination of the channeldepth controlling methods described above. Similar to the probeembodiment, the electrosurgical catheter will preferably include anelectrode array of isolated electrode terminals at it distal end, and areturn electrode proximally recessed from the electrode array.Alternatively, the electrosurgical catheter described above may beguided into the heart and into the desired position using a steerablecatheter body which eliminates the need for a separate steerable guidingcatheter.

In an exemplary embodiment, the electrode array at the distal end of theprobe or catheter is configured such that current flow from at least twoof the electrode terminals is independently controlled based on theelectrical impedance between the electrode terminal and the returnelectrode. Each individual electrode terminal in the electrode array iselectrically connected to a power source which is isolated from each ofthe other electrodes in the array or to circuitry which limits orinterrupts current flow to the electrode when low resistivity material(e.g., blood or electrically conductive saline irrigant) causes a lowerimpedance path between the common electrode and the individual electrodeterminal. The isolated power sources for each individual electrode maybe separate power supply circuits having internal impedancecharacteristics which limit power to the associated electrode terminalwhen a low impedance return path is encountered, may be a single powersource which is connected to each of the electrodes throughindependently actuatable switches or may be provided by independentcurrent limiting elements, such as inductors, capacitors, resistorsand/or combinations thereof, such as resonant circuits. The currentlimiting elements may be provided in the probe, connectors, cable,controller or along the conductive path from the controller to thedistal tip. In addition to the control of power delivery to theelectrodes to effect ablation of the target tissue (e.g., heart wall)while limiting power delivery when low electrical resistivity materialis encountered (e.g., blood), electrosurgical catheter (when employedpercutaneously) may incorporate ultrasonic and/or tissue impedancemeasuring sensor which serve to interrupt power delivery when apreselected channel depth or remaining (uncanalized) wall thickness isreached.

In another aspect of the invention, an electrosurgical system includesthe electrosurgical probe or catheter as described above together withan electrosurgical generator and a delivery mechanism for positioning aradially expandable luminal prothesis into the revascularizing channelsformed by the electrosurgical probe or catheter. The delivery mechanismmay be integral with the electrosurgical instrument, or part of aseparate delivery catheter. The separate delivery catheter usuallyincludes an elongate flexible shaft structure having a proximal end anda distal end. The shaft structure includes a prosthesis receptacle nearthe distal end in or over which a radially compressible tubularprosthesis is carried during maneuvering of the shaft and prosthesiswithin an anatomical lumen. The luminal prostheses will typicallycomprise a resilient, radially compressible, tubular frame having aplurality of openings or slots that allow it to be expanded radiallyoutward into an enlarged configuration. The stent exerts a radial forceagainst the inner channel walls to maintain lumen patency and/ormechanically augment luminal wall strength, thereby maintaining theblood flow from the ventricular cavity to the myocardial tissue.

In yet another aspect of the invention, an instrument guidance system isprovided for detecting an “end point” for each artificial channel and/orfor determining appropriate target sites on the heart wall for formingthe artificial channels. The instrument guidance system will preferablyallow a surgeon to determine when the electrosurgical instrument is nearthe other end of the heart wall (i.e., the outer surface of theepicardium or the inner surface of the endocardium). In the case of thepercutaneous approach in which ablation begins at the endocardium, thedetection system indicates to the surgeon to stop axially translatingthe probe so that the probe does not form a channel completely through aheart wall, which limits bleeding and reduces damage to surroundingtissue structures located at or near the outer surface of the heart. Inaddition, the guidance system will preferably allow the surgeon todetermine an appropriate target site on the heart wall to form thechannel to avoid accidental puncturing of relatively large vessels inthe heart wall. The guidance system may include a fiberoptic viewingsystem or an ultrasound guidance system for determining the targetsites, and/or current limiting circuitry that detects when the probe isadjacent blood vessels and/or the outer or inner edges of the heartwall.

A further understanding of the nature and advantages of the inventionwill become apparent by reference to the remaining portions of thespecification and drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view of the electrosurgical system including anelectrosurgical probe, an electrically conducting liquid supply and anelectrosurgical power supply constructed in accordance with theprinciples of the present invention;

FIG. 2A is an enlarged, cross-sectional view of the distal tip of theelectrosurgical probe of FIG. 1;

FIG. 2B is an end view of the electrosurgical probe of FIG. 1;

FIG. 2C is a cross-sectional view of the proximal end of theelectrosurgical probe of FIG. 2A, illustrating an arrangement forcoupling the probe to the electrically conducting liquid supply of FIG.1;

FIGS. 2D and 2E are cross-sectional views of the distal end of theelectrosurgical probe of FIG. 2A, illustrating one method ofmanufacturing the electrode terminals and insulating matrix of theprobe;

FIG. 3 is a perspective view of the distal tip of anotherelectrosurgical probe that does not incorporate a fluid lumen fordelivering electrically conducting liquid to the target site;

FIG. 4A is an enlarged, cross-sectional view of the distal tip of theelectrosurgical probe of FIG. 3 illustrating an electrode array;

FIG. 4B is an end view of the distal tip of the electrosurgical probe ofFIG. 3;

FIG. 5 is a perspective view of a catheter having a shaft with anelectrosurgical arrangement at its distal end;

FIGS. 6-10A illustrate alternative electrode arrangements for the probesof FIGS. 1-4 or the catheter of FIG. 5;

FIG. 11 is a sectional view of the human heart, illustrating theelectrosurgical catheter of FIG. 5 within the ventricular cavity forperforming a transmyocardial revascularization procedure;

FIG. 12 is a sectional view of the thoracic cavity, illustrating theelectrosurgical probe of FIGS. 3, 4A and 4B in a thoracoscopicrevascularization procedure;

FIG. 13 is a cross-sectional view of the probe of FIGS. 3, 4A and 4Bboring a channel through the myocardium;

FIG. 14 is a cross-sectional view of the probe of FIG. 1 boring achannel through the myocardium;

FIG. 15 depicts an alternative embodiment of the probe of FIG. 14 havingan outer lumen for delivering electrically conductive liquid to thetarget site, and an inner lumen for aspirating fluid and gases from thetransmyocardial channel;

FIG. 16 is a side view of an electrosurgical probe incorporating afiberoptic viewing device and a light generator for sighting the probeonto a target site on the heart tissue;

FIG. 17 schematically illustrates a lumenal prosthesis positioned in arevascularizing channel during a percutaneous procedure to maintainlumen patency;

FIG. 18 is a cross-sectional view of the probe of FIG. 1 boring achannel into the myocardium with an ultrasound tissue thicknessmeasuring device located on a guide catheter;

FIG. 19 is a cross-sectional view of the probe of FIG. 1 boring achannel into the myocardium with an ultrasound tissue thicknessmeasuring device located on an electrosurgical catheter;

FIG. 20 is a cross-sectional view of the probe of FIG. 1 boring achannel into the myocardium with an electrical impedance sensor locatedon an electrosurgical catheter to detect crossing through a surface ofthe heart at a distance L₁ distal to the electrode array;

FIG. 21 is a schematic cross-sectional view of a hemostasis device forsealing artificial revascularizing channels formed by one of theelectrosurgical instruments of the present invention;

FIG. 22 is schematically illustrates a lumenal prosthesis positioned ina revascularizing channel during a thoracoscopic procedure to maintainlumen patency; and

FIG. 23 schematically illustrates a curved revascularizing channelformed by one of the electrosurgical instruments of the presentinvention.

DESCRIPTION OF THE PREFERRED EMBODIMENT

The present invention provides a system and method for selectivelyapplying electrical energy to a target location within or on a patient'sbody. In particular, the present invention provides systems, devices andmethods for increasing the blood flow to the heart by creatingartificial channels or lumens through the myocardium of the heart. Itwill, however, be appreciated that the systems, devices and methods canbe applied equally well to procedures involving other tissues of thebody, as well as to other procedures including open surgery,laparoscopic surgery, thoracoscopic surgery, and other endoscopicsurgical procedures.

The electrosurgical instrument will comprise a shaft having a proximalend and a distal end which supports an active electrode. The shaft mayassume a wide variety of configurations, with the primary purpose beingto mechanically support one or more active electrode and permit thetreating physician to manipulate the electrode(s) from a proximal end ofthe shaft. Usually, an electrosurgical probe shaft will be anarrow-diameter rod or tube, more usually having dimensions which permitit to be introduced into a body cavity, such as the thoracic cavity,through an associated trocar or cannula in a minimally invasiveprocedure, such as arthroscopic, laparoscopic, thoracoscopic, and otherendoscopic procedures. Thus, the probe shaft will typically have alength of at least 5 cm for open procedures and at least 10 cm, moretypically being 20 cm, or longer for endoscopic procedures. The probeshaft will typically have a diameter of at least 1 mm and frequently inthe range from 1 to 10 mm.

The electrosurgical probe may be delivered percutaneously(endoluminally) to the ventricular cavity of the heart by insertionthrough a conventional or specialized guide catheter, or the inventionmay include a catheter having an active electrode array integral withits distal end. The catheter shaft may be rigid or flexible, withflexible shafts optionally being combined with a generally rigidexternal tube for mechanical support. Flexible shafts may be combinedwith pull wires, shape memory actuators, and other known mechanisms foreffecting selective deflection of the distal end of the shaft tofacilitate positioning of the electrode or electrode array. The shaftwill usually include a plurality of wires or other conductive elementsrunning axially therethrough to permit connection of the electrode orelectrode array and the return electrode to a connector at the proximalend of the shaft. Specific shaft designs will be described in detail inconnection with the figures hereinafter.

The present invention may use a single electrode or an electrode arraydistributed over a distal contact surface of the electrosurgicalinstrument. In both configurations, the circumscribed area of theelectrode or electrode array will generally depend on the desireddiameter of the revascularizing channel in the heart. For example,applicant has found that smaller diameter channels tend to remain patentfor a shorter period of time than larger diameter channels. Thus, arelatively large diameter channel (on the order of about 1.5 to 3.0 mm)may be desired to improve lumen patency. The ability to select thediameter of the artificial channels is one of the advantages of thepresent invention over existing LMR procedures, which are typicallylimited by the concentration of light that is required to generatesufficient energy to ablate the tissue during the still or quiescentperiod of the heart (i.e., about 0.08 seconds). Usually, the area of theelectrode array is in the range from 0.25 mm² to 20 mm², preferably from0.5 mm² to 10 mm², and more preferably from about 0.5 mm² to 5.0 mm². Inaddition, the shape of the array and the distal end of the instrumentshaft will also depend on the desired surface area of the channel. Forexample, the ratio of the perimeter of the electrode array to thesurface area may be maximized to increase blood flow from the channel tothe surrounding myocardial tissue. The electrode or electrodes may takethe form of a solid round wire or other solid cross-sectional shapessuch as squares, rectangles, hexagons, triangles, star-shaped or thelike to provide additional edges around the distal perimeter of theelectrodes. Alternatively, the electrode or electrodes may be in theform of hollow metal tubes having a cross-sectional shape which isround, square, hexagonal, rectangular or the like. The envelop oreffective diameter of the individual electrode or electrodes ranges fromabout 0.05 to 3 mm, preferably from about 0.1 to 2 mm.

The electrode array will usually include at least two isolated electrodeterminals, more usually at least four electrode terminals, preferably atleast six electrode terminals, and often 50 or more electrode terminals,disposed over the distal contact surfaces on the shaft. By bringing theelectrode array(s) on the contact surface(s) in close proximity with thetarget tissue and applying high frequency voltage between the array(s)and an additional common or return electrode in direct or indirectcontact with the patient's body, the target tissue is selectivelyablated or cut, permitting selective removal of portions of the targettissue while desirably minimizing the depth of necrosis to surroundingtissue.

As described above, the present invention may use a single activeelectrode or an electrode array distributed over a distal contactsurface of an electrosurgical instrument, such as a probe, a catheter orthe like. The electrode array usually includes a plurality ofindependently current-limited and/or power-controlled electrodeterminals to apply electrical energy selectively to the target tissuewhile limiting the unwanted application of electrical energy to thesurrounding tissue and environment resulting from power dissipation intosurrounding electrically conductive liquids, such as blood, normalsaline, and the like. The electrode terminals may be independentlycurrent-limited by isolating the terminals from each other andconnecting each terminal to a separate power source that is isolatedfrom the other electrode terminals. Alternatively, the electrodeterminals may be connected to each other at either the proximal ordistal ends of the probe to form a single wire that couples to a powersource.

In an exemplary embodiment, each individual electrode terminal in theelectrode array is electrically insulated from all other electrodeterminals in the array within said instrument and is connected to apower source which is isolated from each of the other electrodes in thearray or to circuitry which limits or interrupts current flow to theelectrode when low resistivity material (e.g., blood or electricallyconductive saline irrigant) causes a lower impedance path between thecommon electrode and the individual electrode terminal. The isolatedpower sources for each individual electrode may be separate power supplycircuits having internal impedance characteristics which limit power tothe associated electrode terminal when a low impedance return path isencountered, may be a single power source which is connected to each ofthe electrodes through independently actuatable switches or may beprovided by independent current limiting elements, such as inductors,capacitors, resistors and/or combinations thereof. The current limitingelements may be provided in the probe, connectors, cable, controller oralong the conductive path from the controller to the distal tip. A morecomplete description of a system and method for selectively limitingcurrent to an array of isolated electrode terminals can be found incommonly assigned, co-pending application Ser. No. 08/561,958, filedNov. 22, 1995, the complete disclosure of which has previously beenincorporated herein by reference.

In a preferred aspect, this invention takes advantage of the differencesin electrical resistivity between the target heart tissue and thesurrounding conductive liquid (e.g., isotonic saline irrigant, blood orthe like). By way of example, for any selected level of applied voltage,if the electrical conduction path between the common electrode and oneof the individual electrode terminals within the electrode array isblood (having a relatively low electrical impedance), the currentcontrol means connected to the individual electrode will limit currentflow so that the heating of intervening conductive liquid is minimized.On the other hand, if a portion of or all of the electrical conductionpath between the common or return electrode and one of the individualelectrode terminals within the electrode array is myocardial tissue(having a relatively higher electrical impedance), the current controlcircuitry or switch connected to the individual electrode will allowcurrent flow sufficient for the deposition of electrical energy andassociated ablation or electrical breakdown of the target tissue in theimmediate vicinity of the electrode surface.

It should be clearly understood that the invention is not limited toelectrically isolated electrode terminals, or even to a plurality ofelectrode terminals. For example, the array of active electrodeterminals may be connected to a single lead that extends through theprobe shaft to a power source of high frequency current. Alternatively,the probe may incorporate a single electrode that extends directlythrough the probe shaft or is connected to a single lead that extends tothe power source.

In the case of a single electrode, the invention may also use currentlimiting means to apply electrical energy selectively to the targettissue while limiting the unwanted application of electrical energy tothe surrounding tissue. In this embodiment, the electrode may beconnected to current limiting elements or to circuitry which limits orinterrupts current flow to the electrode when low resistivity material(e.g., blood or electrically conductive saline irrigant) causes a lowerimpedance path between the common electrode and the electrode. Thecurrent limiting elements or circuitry may be configured to completelyinterrupt or modulate current flow to the electrode, for example, when acertain percentage of the electrode surface is in contact with lowresistivity material. In one embodiment, the current flow will bemodulated or completely interrupted when, for example, a large portionof the electrode surface is exposed to fluids and, therefore, not incontact with the target tissue. In this manner, current can beselectively applied to the target tissue, while minimizing current flowto surrounding fluids and adjacent non-target tissue structures.

In addition to the above described methods, the applicant has discoveredanother mechanism for ablating tissue while minimizing the depth ofnecrosis. This mechanism involves applying a high frequency voltagebetween the active electrode surface and the return electrode to develophigh electric field intensities in the vicinity of the target tissuesite. In this embodiment, the active electrode(s) include at least oneactive portion having a surface geometry configured to promotesubstantially high electric field intensities and associated currentdensities between the active portion and the target site when a highfrequency voltage is applied to the electrodes. These high electricfield intensities and current densities are sufficient to break down thetissue by processes including molecular dissociation or disintegration.The high frequency voltage imparts energy to the target site to ablate athin layer of tissue without causing substantial tissue necrosis beyondthe boundary of the thin layer of tissue ablated. This ablative processcan be precisely controlled to effect the volumetric removal of tissuewith minimal heating of or damage to surrounding or underlying tissuestructures.

In an exemplary embodiment, the high electric field intensities at theactive portion of the active electrode(s) may be generated bypositioning the active electrode and target site within an electricallyconducting liquid, such as isotonic saline or other body fluids, such asblood, and applying a high frequency voltage that is sufficient tovaporize the electrically conducting liquid over at least a portion ofthe active electrode in the region between the active portion of theactive electrode and the target tissue. Since the vapor layer orvaporized region has a relatively high electrical impedance, itincreases the voltage differential between the active electrode tip andthe tissue and causes ionization within the vapor layer due to thepresence of an ionizable species (e.g., sodium when isotonic saline isthe electrically conducting fluid). This ionization, under optimalconditions, induces the discharge of energetic electrons and photonsfrom the vapor layer and to the surface of the target tissue. A moredetailed description of this phenomena can be found in application Ser.No. 08/561,958, filed on Nov. 22, 1995, the complete disclosure of whichhas already been incorporated herein by reference.

Suitable electrode surface geometries for producing sufficiently highelectric field intensities to reach the threshold conditions for vaporlayer formation may be obtained by producing sharp edges and/or cornersat the active portion of the active electrode(s). Alternatively, theelectrode(s) may be specifically designed to increase the edge/surfacearea ratio of the active portion through the use of shaped wires (e.g.,square or hexagonal wires) or tubular electrodes offering high electricfield intensities along the inside and outside perimeters of the tubularelectrode. Additionally or alternatively, the active electrodesurface(s) may be modified through chemical, electrochemical or abrasivemethods to create a multiplicity of surface aspirates on the electrodesurface. Suitable electrode designs for use with the present inventionmay be found in co-pending, commonly assigned application Ser. No.08/687,792, filed Jul. 19, 1996, the complete disclosure of which isincorporated herein by reference.

The voltage applied between the common electrode and the electrode arraywill be at high or radio frequency, typically between about 5 kHz and 20MHz, usually being between about 30 kHz and 2.5 MHz, and preferablybeing between about 50 kHz and 1 MHz. The RMS (root mean square) voltageapplied will usually be in the range from about 5 volts to 1000 volts,preferably being in the range from about 50 volts to 800 volts, and morepreferably being in the range from about 60 volts to 500 volts. Thesefrequencies and voltages will result in peak-to-peak voltages andcurrent that are sufficient to vaporize the electrically conductiveliquid and, in turn, create the conditions within the vaporized regionwhich result in high electric fields and emission of energetic photonsand/or electrons to ablate tissue. Typically, the peak-to-peak voltagewill be in the range of 40 to 4000 volts and preferably in the range of100 to 3200 volts and more preferably in the range of 300 to 2400 volts.

As discussed above, the voltage is usually delivered in a waveformhaving a sufficiently high frequency (e.g., on the order of 5 kHz to 20MHz) such that the voltage is effectively applied continuously (ascompared with e.g., lasers claiming small depths of necrosis, which aregenerally delivered in brief pulses at a repetition rate of about 10 to20 Hz). Hence, the duty cycle (i.e., cumulative time in any one-secondinterval that energy is applied) is on the order of about 50% for thepresent invention, as compared with lasers which typically have a dutycycle of about 0.001% to 0.0001%.

With the above voltage and current ranges, applicant has found that theelectrosurgical instrument will usually bore a channel completelythrough the heart wall in about 0.5 to 20.0 seconds, preferably about1.0 to 3.0 seconds, in the continuous mode and preferably about 10 to 15seconds in the pulsed mode. It has been found that channels that areapproximately 0.5 to 3.0 mm in diameter and approximately 1 to 4 cm deepmay be easily and efficiently formed by this method, and that therevascularization procedure dramatically improves the flow of blood tothe heart muscle.

The capability to form the desired channel over a longer period of timesignificantly reduces the amount of instantaneous power required tocomplete the channel. By way of example, CO₂ lasers used for LMRtypically deliver the power for each channel within an elapsed time of0.08 seconds. By contrast, the present invention can be used to completethe canalization of the same sized channel within about 1.0 second. As aresult, the laser requires about 500 to 700 watts to form a 1 mmdiameter channel while the present invention requires 1/12 or about 42to 58 watts to form the same channel. If larger channels are required,the power requirements increase by the square of the ratio of diameters.Hence, to produce a 2 mm channel in 0.08 seconds using a CO₂ laser, therequired power will be four-fold higher or 2000 to 2800 watts whichrequires a very large and very expensive laser. In contrast, the presentinvention can form a 2 mm diameter channel (of the same length as above)in 1 second with an applied power of about 168 to 232 watts.

Usually, the current level will be selectively limited or controlled andthe voltage applied will be independently adjustable, frequently inresponse to the resistance of tissues and/or fluids in the pathwaybetween an individual electrode and the common electrode. Also, theapplied voltage level may be in response to a temperature control meanswhich maintains the target tissue temperature with desired limits at theinterface between the electrode arrays and the target tissue. Thedesired tissue temperature along a propagating surface just beyond theregion of ablation will usually be in the range from about 40° C. to100° C., and more usually from about 50° C. to 60° C. The tissue beingablated (and hence removed from the operation site) immediately adjacentthe electrode array may reach even higher temperatures. A temperaturesensor may be incorporated within the distal end of the electrosurgicaldevice to measure a temperature indicative of the nearby tissue beyondthe ablation boundary.

The preferred power source of the present invention delivers a highfrequency voltage selectable to generate average power levels rangingfrom tens of milliwatts to tens of watts per electrode, depending on thetarget tissue being ablated, the rate of ablation desired or the maximumallowed temperature selected for the probe tip. The power source allowsthe user to select the voltage level according to the specificrequirements of a particular procedure.

The power source may be current limited or otherwise controlled so thatundesired heating of electrically conductive fluids or other lowelectrical resistance media does not occur. In a presently preferredembodiment of the present invention, current limiting inductors areplaced in series with each independent electrode terminal, where theinductance of the inductor is in the range of 10 uH to 50,000 uH,depending on the electrical properties of the target tissue, the desiredablation rate and the operating frequency. Alternatively,capacitor-inductor (LC) circuit structures may be employed, as describedpreviously in co-pending PCT application No. PCT/US94/05168, thecomplete disclosure of which is incorporated herein by reference.Additionally, current limiting resistors may be selected. Preferably,these resistors will have a large positive temperature coefficient ofresistance so that, as the current level begins to rise for anyindividual electrode in contact with a low resistance medium (e.g.,saline irrigant), the resistance of the current limiting resistorincreases significantly, thereby minimizing the power delivery from saidelectrode into the low resistance medium (e.g., saline irrigant).

As an alternative to such passive circuit structures, regulated currentflow to each electrode terminal may be provided by a multi-channel powersupply. An applied voltage with active current sensing circuitry isprovided for each individual electrode terminal to control currentwithin a range which will limit power delivery through a low resistancepath, e.g., isotonic saline irrigant, and would be selected by the userto achieve the desired rate of cutting or ablation. Such a multi-channelpower supply thus provides a voltage source with controlled currentlimits with selectable voltage level in series with each electrodeterminal, wherein all electrodes will operate at or below the same, userselectable maximum current level. Current flow to all electrodeterminals could be periodically sensed and stopped if the temperaturemeasured at the surface of the electrode array exceeds user selectedlimits. Particular control system designs for implementing this strategyare well within the skill of the art.

Yet another alternative involves the use of one or several powersupplies which allow one or several electrodes to be simultaneouslyenergized and which include active control means for limiting currentlevels below a preselected maximum level. In this arrangement, only oneor several electrodes would be simultaneously energized for a briefperiod. Switching means would allow the next one or several electrodesto be energized for a brief period. By sequentially energizing one orseveral electrodes, the interaction between adjacent electrodes can beminimized (for the case of energizing several electrode positioned atthe maximum possible spacing within the overall envelope of theelectrode array) or eliminated (for the case of energizing only a singleelectrode at any one time). As before, a resistance measurement meansmay be employed for each electrode prior to the application of powerwherein a (measured) low resistance (below some preselected level) willprevent that electrode from being energized during a given cycle.

It should be clearly understood that the invention is not limited toelectrically isolated electrode terminals, or even to a plurality ofelectrode terminals. For example, the array of active electrodeterminals may be connected to a single lead that extends through theprobe shaft to a power source of high frequency current. Alternatively,the probe may incorporate a single electrode that extends directlythrough the probe shaft or is connected to a single lead that extends tothe power source.

In yet another aspect of the invention, the control system is “tuned” sothat it will not apply excessive power to the blood (e.g., in the leftventricle), once it crosses the wall of the heart and enters the chamberof the left ventricle. This minimizes the formation of a thrombus in theheart (i.e., will not induce thermal coagulation of the blood). Thecontrol system may include an active or passive architecture, and willtypically include a mechanism for sensing resistance between a pair(s)of active electrodes at the distal tip, or between one or more activeelectrodes and a return electrode, to sense when the electrode array hasentered into the blood-filled chamber of the left ventricle.Alternatively, current limiting means may be provided to preventsufficient joulean heating in the lower resistivity blood to causethermal coagulation of the blood. In another alternative embodiment, anultrasound transducer at the tip of the probe can be used to detect theboundary between the wall of the heart and the blood filled leftventricle chamber, turning off the electrode array just as the probecrosses the boundary.

Referring to the drawings in detail, wherein like numerals indicate likeelements, an electrosurgical system 11 is shown in FIG. 1 constructedaccording to the principles of the present invention. Electrosurgicalsystem 11 generally comprises an electrosurgical instrument or probe orcatheter 10 connected to a power supply 28 for providing high frequencyvoltage to electrosurgical instrument 10 and a liquid source 21 providedfor supplying electrically conducting fluid 50 to probe 10.

In an exemplary embodiment as shown in FIG. 1, electrosurgical probe 10includes an elongated shaft 13 which may be flexible or rigid, withflexible shafts optionally including support cannulas or otherstructures (not shown). It will be recognized that the probe shown inFIG. 1 will generally be employed in open or thoracoscopic proceduresthrough intercostal penetrations in the patient. For endoluminalprocedures into the ventricle, a delivery catheter 200 (FIGS. 6 and 11)will typically be employed, as discussed below. Probe 10 includes aconnector 19 at its proximal end and an array 12 of electrode terminals58 disposed on the distal tip of shaft 13. A connecting cable 34 has ahandle 22 with a connector 20 which can be removably connected toconnector 19 of probe 10. The proximal portion of cable 34 has aconnector 26 to removably couple probe 10 to power supply 28. Theelectrode terminals 58 are electrically isolated from each other andeach of the terminals 58 is connected to an active or passive controlnetwork within power supply 28 by means of a plurality of individuallyinsulated conductors 42 (see FIG. 2A). Power supply 28 has a selectionmeans 30 to change the applied voltage level. Power supply 28 alsoincludes means for energizing the electrodes 58 of probe 10 through thedepression of a pedal 39 in a foot pedal 37 positioned close to theuser. The foot pedal 37 may also include a second pedal (not shown) forremotely adjusting the voltage level applied to electrodes 58. Thespecific design of a power supply which may be used with theelectrosurgical probe of the present invention is described in parentapplication PCT US 94/051168, the full disclosure of which haspreviously been incorporated herein by reference.

Referring to FIGS. 2A and 2B, the electrically isolated electrodeterminals 58 are spaced-apart over an electrode array surface 82. Theelectrode array surface 82 and individual electrode terminals 58 willusually have dimensions within the ranges set forth above. In thepreferred embodiment, the electrode array surface 82 has a circularcross-sectional shape with a diameter D (FIG. 2B) in the range from 0.3mm to 4 mm. Electrode array surface 82 may also have an oval orrectangular shape, having a length L in the range of 1 mm to 20 mm and awidth W in the range from 0.3 mm to 7 mm, as shown in FIG. 8 (discussedbelow). The individual electrode terminals 58 will protrude over theelectrode array surface 82 by a distance (H) from 0 mm to 2 mm,preferably from 0 mm to 1 mm (see FIG. 2A).

The electrode terminals 58 are preferably composed of a electricallyconductive metal or alloy, such as platinum, titanium, tantalum,tungsten, niobium, stainless steel, and the like. A preferred materialfor terminals 58 is tungsten because of its known biocompatibility andresistance to erosion under the application of high voltages. As shownin FIG. 2B, the electrode terminals 58 are anchored in a support matrix48 of suitable insulating material (e.g., ceramic, glass/ceramic, orglass material, such as alumina, silica glass and the like) which couldbe formed at the time of manufacture in a flat, hemispherical or othershape according to the requirements of a particular procedure. In anexemplary embodiment, the support matrix 48 will comprise an inorganicinsulator, such as ceramic, glass, glass/ceramic or a high resistivitymaterial, such as silicon or the like. An inorganic material isgenerally preferred for the construction of the support matrix 48 sinceorganic or silicone based polymers are known to rapidly erode duringsustained periods of the application of high voltages between electrodeterminals 58 and the return electrode 56 during tissue ablation.However, for situations in which the total cumulative time of appliedpower is less than about one minute, organic or silicone based polymersmay be used without significant erosion and loss of material of thesupport matrix 48 and, therefore, without significant reduction inablation performance.

As shown in FIG. 2A, the support matrix 48 is adhesively joined tosupport member 9, which extends most or all of the distance betweenmatrix 48 and the proximal end of probe 10. In a particularly preferredconstruction technique, support matrix 48 comprises a plurality of glassor ceramic hollow tubes 400 (FIG. 2D) extending from the distal end ofshaft 13. In this embodiment, electrode terminals 58 are each insertedinto the front end of one of the hollow tubes 400 and adhered to thehollow tubes 400 so that the terminals 58 extend distally from eachhollow tube 400 by the desired distance, H. The terminals 58 arepreferably bonded to the hollow tubes 400 by a sealing material 402(e.g., epoxy) selected to provide effective electrical insulation, andgood adhesion to both the hollow tubes 400 and the electrode terminals58. Alternatively, hollow tubes 400 may be comprised of a glass having acoefficient of thermal expansion similar to that of electrode terminal58 and may be sealed around the electrode terminal 58 by raising thetemperature of the glass tube to its softening point according to theprocedures commonly used to manufacture glass-to-metal seals. Referringto FIG. 2D, lead wires 406, such as insulation 408 covered copper wires,are inserted through the back end of the hollow tubes 400 and coupled tothe terminals 58 with a suitable conductive adhesive 404. The glasstube/electrode terminal assembly is then placed into the distal end ofsupport member 9 to form the electrode array as shown in FIG. 2E.Alternatively, the lead wire 406 and electrode terminal 58 may beconstructed of a single wire (e.g., stainless steel or nickel alloy)with insulation 408 removed over the length of the wire inserted intothe hollow tube 400. As before, sealing material 402 is used to sealannular gaps between hollow tube 400 and electrode terminal 58 and toadhesively join electrode terminal 58 to hollow tube 400. Other featuresof construction are discussed above and shown in FIG. 2E.

In the embodiment shown in FIGS. 2A and 2B, probe 10 includes a returnelectrode 56 for completing the current path between electrode terminals58 and power supply 28. Shaft 13 preferably comprises an electricallyconducting material, usually metal, which is selected from the groupconsisting of stainless steel alloys, platinum or its alloys, titaniumor its alloys, molybdenum or its alloys, and nickel or its alloys. Thereturn electrode 56 may be composed of the same metal or alloy whichforms the electrode terminals 58 to minimize any potential for corrosionor the generation of electrochemical potentials due to the presence ofdissimilar metals contained within an electrically conductive fluid 50,such as isotonic saline (discussed in greater detail below).

As shown in FIGS. 2A, 2B and 2C, return electrode 56 extends from theproximal end of probe 10, where it is suitably connected to power supply28 via connectors 19, 20, to a point slightly proximal of electrodearray surface 82, typically about 0.5 to 10 mm and more preferably about1 to 10 mm. Shaft 13 is disposed within an electrically insulativejacket 18, which is typically formed as one or more electricallyinsulative sheaths or coatings, such as polyester,polytetrafluoroethylene, polyimide, and the like. The provision of theelectrically insulative jacket 18 over shaft 13 prevents directelectrical contact between shaft 13 and any adjacent body structure orthe surgeon. Such direct electrical contact between a body structure andan exposed return electrode 56 could result in unwanted heating andnecrosis of the structure at the point of contact causing necrosis.

In the embodiment shown in FIGS. 2A-2C, return electrode 56 is notdirectly connected to electrode terminals 58. To complete this currentpath so that terminals 58 are electrically connected to return electrode56 via target tissue 52, electrically conducting liquid 50 (e.g.,isotonic saline) is caused to flow along liquid paths 83. A liquid path83 is formed by annular gap 54 between outer return electrode 56 andtubular support member 78. An additional liquid path 83 may be formedbetween an optional inner lumen 57 within an inner tubular member 59.The electrically conducting liquid 50 flowing through fluid paths 83provides a pathway for electrical current flow between target tissue 52and return electrode 56, as illustrated by the current flux lines 60 inFIG. 2A. When a voltage difference is applied between electrode array 12and return electrode 56, high electric field intensities will begenerated at the distal tips of terminals 58 with current flow fromarray 12 through the target tissue to the return electrode, the highelectric field intensities causing ablation of tissue 52 in zone 88.

FIG. 2C illustrates the proximal or connector end 70 of probe 10 in theembodiment of FIGS. 2A and 2B. Connector 19 comprises a plurality ofindividual connector pins 74 positioned within a housing 72 at theproximal end 70 of probe 10. Electrode terminals 58 and the attachedinsulating conductors 42 extend proximally to connector pins 74 inconnector housing 72. Return electrode 56 extends into housing 72, whereit bends radially outward to exit probe 10. As shown in FIG. 1, a liquidsupply tube 15 removably couples liquid source 21, (e.g., a bag ofelectrically conductive fluid elevated above the surgical site or havinga pumping device), with return electrode 56. Preferably, an insulatingjacket 14 covers the exposed portions of electrode 56. One of theconnector pins 76 is electrically connected to return electrode 56 tocouple electrode 56 to power supply 28 via cable 34. A manual controlvalve 17 may also be provided between the proximal end of electrode 56and supply tube 15 to allow the surgical seam to regulate the flow ofelectrically conducting liquid 50.

FIGS. 3, 4A and 4B illustrate another preferred embodiment of thepresent invention. In this embodiment, the probe does not include afluid channel for directing electrically conducting liquid to the targetsite. Applicant has found that the fluids in the patient's heart tissue,such as blood, usually have a sufficient amount of electricalconductivity to complete the electrical path between the activeelectrode array and the return electrodes. In addition, these fluidswill often have the requisite properties discussed above forestablishing a vapor layer, creating regions of high electric fieldsaround the edges of electrode terminals 58 and inducing the discharge ofenergetic electrons and photons from the vapor layer to the surface ofthe target tissue to effect ablation.

As shown in FIG. 3, electrosurgical probe 100 has a shaft 102 with anexposed distal end 104 and a proximal end (not shown) similar to theproximal end shown in FIG. 2C. Aside from exposed distal end 104, whichfunctions as the return electrode in this embodiment, the entire shaft102 is preferably covered with an electrically insulative jacket 106,which is typically formed as one or more electrically insulative sheathsor coatings, such as polyester, polytetrafluoroethylene, polyimide, andthe like, to prevent direct electrical contact between shaft 102 and anyadjacent body structure or the surgeon. Similar to the previousembodiment, probe 100 includes an array 108 of active electrodeterminals 110 having substantially the same applied potential. Terminals110 extend from the an insulating inorganic matrix 112 attached todistal end 104 of shaft 102. As discussed above, matrix 112 is onlyshown schematically in the drawings, and preferably comprises an arrayof glass or ceramic tubes extending from distal end 104 or is a ceramicspacer through which electrode terminals 110 extend.

The electrode array may have a variety of different configurations otherthan the one shown in FIGS. 3, 4A and 4B. For example, as shown in FIG.8, the distal end of the shaft 102 and/or the insulating matrix may havea substantially rectangular shape with rounded corners so as to maximizethe perimeter length to cross-sectional area ratio of the distal tip ofthe probe. As shown, electrode array surface 120 has a rectangular shapehaving a width in the range of 2 mm to 5 mm and a height in the range of1 mm to 2 mm. Increasing the perimeter of the artificial channel mayhave advantages in revascularizing the heart because the blood flowingthrough the artificial channel will have a greater area to pass into theheart tissue for a given cross-sectional area. Thus, this configurationis a more efficient method of increasing blood flow within themyocardium.

In another embodiment, the return electrode is positioned on the frontor distal face of the probe. This configuration inhibits current flowwithin the tissue on the sides of probe as it forms the revascularizingchannel. In one configuration, for example (shown in FIG. 9), theelectrode array surface 122 includes multiple pairs of electrodes, witheach pair of electrodes including an active electrode 126 and a returnelectrode 128. Thus, as shown, the high frequency current passes betweenthe pairs across the distal surface 122 of the probe. In anotherconfiguration (shown in FIG. 10), the return or common electrode 130 ispositioned in the center of the distal probe surface 132 and the activeelectrodes 134 are positioned at its perimeter. In this embodiment, theelectrosurgical current will flow between active electrodes 134 at theperimeter of distal surface 132 and return electrode 130 at its centerto form the revascularizing channel. This allows the surgeon to moreprecisely control the diameter of the revascularizing channel becausethe current will generally flow radially the outer electrodes 134 andthe return electrode 130. For this reason, electrodes 134 willpreferably be positioned on the perimeter of distal surface 132 (i.e.,further radially outward than shown in FIG. 10A) to avoid tearing ofnon-ablated heart tissue by the perimeter of the probe shaft.

FIG. 6 illustrates yet another embodiment of an electrosurgical probe100′ according to the present invention. In this embodiment, the distaltip of the probe has a conical shape and includes an array of activeelectrodes along the conical surface 140. A conical shape provides lessresistance to the advancement of the probe through dense tissue. Asshown in FIG. 6, insulating matrix 142 tapers in the distal direction toform conical distal surface 140. The electrode array 144 extends fromdistal surface 140, with each electrode terminal 146 arranged toprotrude axially from the conical surface 140 (i.e., rather thanprotruding perpendicularly from the surface 140). With thisconfiguration, the electrodes 146 do not extend radially outward fromthe conical surface 140, which reduces the risk of electric currentflowing radially outward to heart tissue surrounding the revascularizingchannel. In addition, the high electric field gradients generated by theelectric current concentrate near the active electrode surfaces andtaper further away from these surfaces. Therefore, this configurationplaces these high electric field gradients within the diameter of thedesired channel to improve ablation of the channel, while minimizingablation of tissue outside of the desired channel.

FIG. 5 illustrates a preferred delivery catheter 200 for introducing anelectrosurgical probe 202 through a percutaneous penetration in thepatient, and endoluminally delivering probe 202 into the ventricle ofthe heart (this method described in detail below). Catheter 200generally includes a shaft 206 having a proximal end 208 and a distalend 210. Catheter 200 includes a handle 204 secured to proximal end 208of shaft, and preferably a deflectable tip 212 coupled to distal end 210of shaft 206. Probe 202 extends from proximal end, preferably by adistance of about 100 to 200 cm. Handle 204 includes a variety ofactuation mechanisms for manipulating tip 212 within the patient'sheart, such as a tip actuation slide 214 and a torque ring 216, as wellas an electrical connector 218. Catheter shaft 206 will generally defineone or more inner lumens 220, and one or more manipulator wires andelectrical connections (not shown) extending through the lumens to probe202.

With reference to FIGS. 11-22, methods for increasing the blood flow tothe heart through a transmyocardial revascularization procedure to formartificial channels through the heart wall to perfuse the myocardiumwill now be described. This procedure is an alternative to coronaryartery bypass surgery for treating coronary artery disease. The channelsallow oxygen enriched blood flowing into the ventricular cavity todirectly flow into the myocardium rather than exiting the heart and thenflowing back into the myocardium through the coronary arteries.

As shown in FIG. 11, electrosurgical probe 202 is positioned into theleft ventricular cavity 258 of the heart. Electrosurgical probe 202 maybe introduced into the left ventricle 250 in a variety of proceduresthat are well known in the art, such as a percutaneous, minimallyinvasive procedures. In the representative embodiment, probe 202 isintroduced into the vasculature of the patient through a percutaneouspenetration 360 and axially translated via delivery catheter 200 throughone of the major vessels, such as the femoral artery 346, through theaorta 254 to the left ventricular cavity 258. A viewing scope (notshown) may also be introduced through a percutaneous position to aposition suitable for viewing the target location in the left ventricle258.

Once positioned within the patient's ventricle 258, probe 202 is alignedwith the heart wall 260 to form one or more artificial channels 264 forincreasing blood flow to the myocardium 262. As shown in FIG. 14, thechannels 264 will preferably extend from the endocardium 266 a desireddistance through the myocardium 262 without perforating the exterior ofthe epicardium 268 to inhibit blood from flowing into the thoraciccavity. Preferably, the surgeon will employ a detection or instrumentguidance system 350, (discussed below in reference to FIGS. 16, 18, 19and 20) on probe 202, or another instrument, to determine when the probeis near the outer surface of the epicardium 268. The location ofchannels 264 may be selected based on familiar endocardial anatomiclandmarks. Alternatively, instrument guide system 350 may be used toselect target sites on the heart wall, as discussed below.

As shown in FIG. 14, guide catheter 200 is positioned adjacent the innerendocardial wall and probe 202 is axially translated so that the activeelectrode 270 at its distal end is positioned proximate the hearttissue. In this embodiment, the probe 202 includes a single, annularelectrode 270 at its distal tip for ablation of the heart tissue.However, it will be readily recognized that the probe may include anarray of electrode terminals as described in detail above. While viewingthe region with an endoscope (not shown), voltage can be applied frompower supply 28 (see FIG. 1) between active electrode 270 and annularreturn electrode 272. The boring of channel 264 is achieved by engagingactive electrode 270 against the heart tissue or positioning activeelectrode 270 in close proximity to the heart tissue whilesimultaneously applying voltage from power supply 28 and axiallydisplacing probe 202 through channel 264. To complete the current pathbetween the active and return electrodes 270, 272, electricallyconducting irrigant (e.g., isotonic saline) will preferably be deliveredfrom liquid supply 21 through annular liquid path 274 between returnelectrode 272 and tubular shaft 200 to the target site. Alternatively,the site may already be submerged in liquid, or the liquid may bedelivered through another instrument. The electrically conducting liquidprovides a pathway for electrical current flow between the heart tissueand return electrode 272, as illustrated by the current flux lines 278in FIG. 15.

FIG. 15 illustrates an alternative embodiment of the probe of FIG. 14.In this embodiment, the probe 280 includes a central lumen 282 having aproximal end attached to a suitable vacuum source (not shown) and anopen distal end 286 for aspirating the target site. To complete thecurrent path between the active electrode 270 and return electrode 272,electrically conducting irrigant (e.g., isotonic saline) will preferablybe delivered from liquid supply 21 (shown in FIG. 1) through annularliquid path 274 between return electrode 272 and tubular shaft 200 tothe target site. The active electrode is preferably a single annularelectrode 288 surrounding the open distal end 286 of central lumen 282.Central lumen 282 is utilized to remove the ablation products (e.g.,liquids and gases) generated at the target site and excess electricallyconductive irrigant during the procedure.

An alternative embodiment of the percutaneous, endocardial canalizationapproach is shown in FIG. 23. In this embodiment, electrosurgicalcatheter 100 can be guided by the surgeon or surgical assistant duringthe canalization of channel 264 using external handpiece 340 shown inFIG. 11. In this embodiment, the distal portion of the electrosurgicalcatheter 100 can be caused to follow a curved path to effect a curvedartificial channel 264′. By forming a curved artificial channel 264′,the total surface area of the artificial channel can be extended so thatsaid channel is longer than the total thickness L₉ of the heart wall260. In addition, by forming a curved artificial channel 264′ of propercurvature as shown in FIG. 23, the penetration of the epicardium 268 canbe avoided. Still further, the curved artificial channel 264′ can becontinued forming a complete “U” shaped channel which reenters theventricular cavity 258 providing one continuous channel which penetratesthe endocardium at two locations but does not penetrate through theepicardium 268.

FIG. 12 illustrates a thoracoscopic procedure for revascularizing themyocardium from the outer wall or epicardium 268 inward to theendocardium 266. At least one intercostal penetration is made in thepatient for introduction of electrosurgical probe 100 (FIG. 3) into thethoracic cavity 302. The term “intercostal penetration” as used hereinrefers to any penetration, in the form of a small cut, incision, hole orcannula, trocar sleeve or the like, through the chest wall between twoadjacent ribs which does not require cutting, removing, or significantlydisplacing or retracting the ribs or sternum. Usually, the intercostalpenetration will require a puncture or incision of less than about 5 cmin length. Preferably, the intercostal penetration will be a trocarsleeve 300 having a length in the range from about 2-15 cm, and aninternal diameter in the range from 1 to 15 mm, commonly known asthoracic trocars. Suitable trocar sleeves are available from UnitedStates Surgical Corp. of Norwalk, Conn., under the brand name“Thoracoport”™. A viewing scope (not shown) may also be introducedthrough a trocar sleeve to a position suitable for viewing the targetlocation on the heart wall 260. A viewing scope (not shown) may also beintroduced through the same or another intercostal penetration into thethoracic cavity 302 to a position suitable for viewing the targetlocation 360 on the surface of the epicardium 268 of the heart. Theviewing scope can be a conventional laparoscope or thoracoscope, whichtypically comprise a rigid elongated tube containing a lens system andan eyepiece or camera mount at the proximal end of the tube. A smallvideo camera is preferably attached to the camera mount and connected toa video monitor to provide a video image of the procedure. This type ofscope is commercially available from Baxter Healthcare Corporation ofDeerfield, Ill. or United States Surgical Corporation of Norwalk, Conn.

As shown in FIG. 13, one or more artificial channels 264 are formed bythe electrosurgical probe 100 from the outer wall or epicardium 268through the myocardium 262 and the inner wall or endocardium 266 intothe ventricular cavity 258. Similar to the above described method,electrode array 108 is positioned in close proximity to the heart tissuewhile simultaneously applying voltage from power supply 28 and axiallydisplacing probe 100 through channel 264. In this embodiment, however,electrically conducting liquid is not supplied to the target site tocomplete the current path between the active electrode terminals 110 andreturn electrode 102. Applicant has found that the fluids in thepatient's heart tissue, such as blood, usually have a sufficient amountof electrical conductivity to complete the electrical path between theactive electrode array and the return electrodes. In addition, thesefluids will often have the requisite properties discussed above forestablishing a vapor layer and inducing the discharge of energeticelectrons and photons from the vapor layer to the surface of the targettissue as well as the creation of high electric fields to effect theablation of tissue.

To inhibit blood from flowing through channels 264 into the thoraciccavity, the channels 264 will preferably be sealed at the epicardium 268as soon as possible after they have been formed. One method for sealingthe artificial channel 264 at the epicardium 268 is to insert a collagenhemostasis device 480 (shown in FIG. 21) using a trocar 300, a cannula484 and a syringe-like delivery system 486. The collagen, unaffected byantiplatelet or anticoagulant agents that may be present in thepatient's blood stream, attracts and activates platelets from the blood482, rapidly forming a “glue”-like plug near the surface of theepicardium 268 of the newly formed channel 264. Suitable collagenhemostasis devices are available from Datascope Corporation, Montval,N.J. under the brand name “VasoSeal™”. The deployment of the collagenhemostasis device 480 is accomplished with the aid of a viewing scope(not shown) which may also be introduced through a trocar sleeve to aposition suitable for viewing the target location on the heart wall 260.

To facilitate this sealing procedure, the electrosurgical probe 354 willpreferably include a guidance system 350 (FIG. 16) for determining whenthe probe is close to the inner surface of the endocardium 266 so thatthe surgeon can prepare to withdraw probe 100 and seal the channel 264.

In both of the above embodiments, the present invention provideslocalized ablation or disintegration of heart tissue to form arevascularization channel 264 of controlled diameter and depth. Usually,the diameter will be in the range of 0.5 mm to 3 mm, preferably betweenabout 1 to 2 mm. Preferably, the radio frequency voltage will be in therange of 300 to 2400 volts peak-to-peak to provide controlled rates oftissue ablation and hemostasis while minimizing the depth of necrosis oftissue surrounding the desired channel. This voltage will typically beapplied continuously throughout the procedure until the desired lengthof the channel 264 is completely formed. However, the heartbeat may bemonitored and the voltage applied in pulses that are suitably timed withthe contractions (systole) of the heart.

Ablation of the tissue may be facilitated by axially reciprocatingand/or rotating the electrosurgical probe a distance of between about 1to 5 mm. This axial reciprocation or rotation allows the electricallyconducting liquid (FIG. 14) to flow over the tissue surface beingcanalized, thereby cooling this tissue and preventing significantthermal damage to the surrounding tissue cells.

FIG. 16 illustrates one representative guidance system 350 for guidingan electrosurgical probe 354 to target sites on the heart wall. Guidancesystem 350 is provided for detecting an “end point” for each artificialchannel and/or for determining appropriate target sites on the heartwall for forming the artificial channels. The instrument guidance system350 will preferably allow a surgeon to determine when theelectrosurgical probe 354 (or an electrosurgical catheter) is near theother end of the heart wall (i.e., the outer edge of the epicardium orthe inner edge of the endocardium). The guidance system 350 indicates tothe surgeon to stop axially translating the probe 354 so that the probedoes not form a channel completely through a heart wall, which limitsbleeding and/or reduces damage to surrounding tissue structures orblood. Alternatively or in addition, the guidance system 354 will allowthe surgeon to determine an appropriate target site 360 on the heartwall to form the channel to avoid accidental puncturing of relativelylarge vessels in the heart wall.

In one embodiment shown in FIG. 16, the instrument guidance system 350includes a fiberoptic viewing cable 356 within electrosurgical probe354, and a visible light generator 358, such as a laser beam, integralwith the probe for illuminating a target site 360 on the heart wall.Note that both light generator 358 and fiberoptic cable 356 may becoupled to an instrument other than probe 354. The fiberoptic viewingcable 356 sites the target site 360 illuminated by the visible lightgenerator 358 to locate where the probe 354 will bore the hole. Thisallows the surgeon to avoid puncturing larger blood vessels 362 on theheart wall (e.g., coronary arteries or veins). Visible light generator358 may also be used to determine when the distal end 364 of probe 354is close to the opposite side of the heart wall, or when the probe 354has completely penetrated through the heart wall into the ventricularcavity 258.

In a second embodiment, the detection system is an ultrasound guidancesystem that transmits sound waves onto the heart wall to facilitatecanalization of the heart.

Referring to FIGS. 18 and 19, an ultrasound tissue thickness measuringsystem may be incorporated within the electrosurgical catheter 100 orguide tube 200 to measure the thickness of the heart wall 260 adjacentto the distal probe tip 270 and thereby allow the surgeon to pre-set thedepth of each channel using adjustable stop 352 on handpiece 340 (FIG.11) before energizing the electrosurgical catheter 100 and ablating theheart tissue. In the embodiment shown in FIG. 18, an ultrasonictransducer 310 affixed to the distal end of guide tube 200 and connectedto an external ultrasonic generator and sensing system (not shown) vialead 312, transmits pulses of ultrasound into the heart tissue in theform of emitted ultrasound signal 314 and the ultrasound generator andsensing system measures the delay time for reflected ultrasound signal316 to return from the boundary of the heart wall at the surface ofepicardium 268. This delay time can be translated into a thickness ofthe entire heart wall and allow the surgeon to adjust the maximum traveldistance of electrosurgical catheter 100 using mechanical stop 352 (FIG.11) to prevent the length of channel 264 from extending through theouter surface of the epicardium 268. The surgeon can choose to stop thecanalization of the heart at any selected distance of the epicardiumwhich may typically be in the range from about 1 mm to 10 mm.

A third embodiment is shown in FIG. 19 wherein an ultrasonic transducer310 is affixed to the distal end of electrosurgical catheter 100 andconnected to an external ultrasonic generator and sensing system (notshown) via leads 312, and transmits pulses of ultrasound into the hearttissue in the form of emitted ultrasound signal 314. The ultrasoundgenerator and sensing system measures the delay time for reflectedultrasound signal 316 to return from the boundary of the heart wall atthe surface of epicardium 268. This measured delay time can betranslated into the distance between the distal tip 270 ofelectrosurgical catheter 100 and the surface of the epicardium 268. Inthis arrangement, the surgeon can observe where the channel 264 reachesthe preferred distance from the epicardium 268 and interrupt theapplication of power and advancement of electrosurgical catheter 100.Alternatively, the preferred minimum thickness of the uncanalized heartwall 260 (i.e., the minimum distance from the bottom of channel 264 tothe surface of the epicardium 268) can be preselected by the surgeon.When this distance is reached based on the thickness of the uncanalizedheart wall measured using the ultrasonic generator and sensor system(now shown), the ultrasonic generator and sensor system provides anelectrical signal to the power source for the electrosurgical catheter100 to interrupt the applied voltage, thereby ending the canalizationprocess and limiting the depth of channel formed. In this manner, thesurgeon may hear an audible tone and will “feel” the catheteradvancement stop at the moment the applied voltage is interrupted.

A fourth embodiment is shown in FIG. 20 in which electrosurgicalcatheter 100 includes a small diameter tissue electrical impedancemeasurement sensor 319 which extends distal to the tissue ablatingelectrode array 270 by a distance L₆. The impedance measurement sensor319 detects the outer surface of the epicardium 268 as it enters aregion of different electrical impedance (viz, the fluid-filled cavitysurrounding the heart). In the present embodiment, sensor tip 320 mayinclude a first impedance measurement electrode 321 and a secondimpedance measurement electrode 323. A small, high-frequency potentialis applied between first and second impedance measurement electrodes 321and 323 causing current flow between first and second impedancemeasurement electrodes 321 and 323 as indicated by current flux lines322. As the first and second electrodes 321 and 323 emerge from theepicardium 268 into cavity 318 surrounding the heart, the change in theelectrical impedance is measured and may be indicated by an audiblesignal and/or may be used as a direct feedback control signal tointerrupt the application of voltage to the electrosurgical catheter 100by generator 28 (FIG. 1). By this method, the forward advancement of theelectrosurgical catheter 100 can be limited to a preselected distance L₆between the bottom of channel 264 and the surface of the epicardium 268.

In a fifth embodiment shown in FIG. 13, the guidance system utilizesimpedance measurement circuitry integrated with the ablating electrodes110 to detect when the electrosurgical catheter probe 100 is adjacentblood vessels and/or the outer or inner boundaries of the heart wall.Specifically, the current limiting circuitry includes a number ofimpedance monitors coupled to each electrode terminal to determine theimpedance between the individual electrode terminal 110 and the returnof common electrode 102. Thus, for example, if the measured impedancesuddenly decreases at electrode terminals 110 at the tip of the probe100, the applied voltage will be interrupted to avoid power delivery toblood filled ventricular cavity 258 of the heart, thereby avoidingformation of a thrombus or damage to other tissue structures within theventricular cavity 258.

FIG. 17 illustrates a method for implanting a luminal prosthesis, suchas a stent or stent-graft 370, into the artificial channels 264 formedby one of the electrosurgical probes or catheters of the presentinvention to maintain the patency of these channels 264. The stents 370are usually compressed into a narrow-diameter configuration (not shown),and advanced endoluminally to one of the artificial channels 264 in theheart wall with a conventional or specialized delivery catheter (notshown). Alternatively, the electrosurgical probe may be designed todelivery and implant the stents 370 at the target site. The stents 370will typically comprise a resilient, radially compressible, tubularframe 372 having a proximal end 374, a distal end 376, and an axiallumen 380 therebetween. The tubular frame 372 includes a plurality ofopenings or slots (not shown) that allow it to be expanded radiallyoutward into the enlarged configuration shown in FIG. 17 by conventionalmethods, such as shape memory alloys, expandable balloons, and the like.The stent 370 exerts a radial force against the inner channel walls 382to maintain lumen patency and/or mechanically augment luminal wallstrength, thereby maintaining the blood flow from the ventricular cavityto the myocardium. The stent 370 may also include a graft or liner 384for inhibiting cell proliferation and occlusion through the openings andslots of frame 372.

In a first embodiment shown in FIG. 17, the stent 370 is introduced intothe artificial channel 264 during a percutaneous procedure asillustrated in FIG. 11. In this embodiment, the length of each channel264 and hence the length of each stent 370 extends only partiallythrough the entire thickness of heart wall 260.

In a second embodiment shown in FIG. 22, the stent is introduced intothe artificial channel 264 during a thoracoscopic procedure asillustrated in FIG. 12. In this embodiment, the length of eachartificial channel 264 extends completely through the heart wall 260 toallow the blood within the ventricular cavity 258 to circulate withinthe majority of the length of the artificial channel 264. However, inthis embodiment, the stent 370 is placed in the distal portion of theartificial channel 264 as shown in FIG. 22 extending to the endocardium266 to maintain patency of the artificial channel 264 over length L₇ ofheart wall 260. Following insertion and deployment of stent 370, theproximal portion of artificial channel 264 may be sealed using collagenhemostasis device 480, or the like, as described hereinbefore related toFIG. 21 using a cannula 484 and syringe-like delivery system 486 asshown in FIG. 21. The collagen hemostasis device 480 attracts andactivates platelets from the blood 482, rapidly forming a “glue”-likeplug near the surface of the epicardium 268 of the newly formedartificial channel 264. Alternatively, a collagen hemostasis device maybe deployed through a central lumen 59 integral with the electrosurgicalprobe or catheter as illustrated in FIG. 2A. The collagen hemostasisdevice can be compressed to fit within a lumen 59 whose diameter issmaller than that of the artificial channel 264. When ejected form theconfining lumen 59, the collagen hemostasis device 480 expands to fillthe full diameter of the artificial channel 264 over length L₈ as shownin FIG. 22. Also, such a system for the deployment of a collagenhemostasis device 480 or the like may be integrated with theelectrosurgical catheter 100 used for the percutaneous canalization ofartificial channels 264 according to method illustrated in FIG. 11.Referring to FIGS. 2A, 11 and 22, in a percutaneous approach using thissealing method, the artificial channel 264 would be formed through theentire thickness of the heart wall 260. Once the surface of theepicardium 268 is penetrated, the position of the tip 202 of theelectrosurgical catheter 100 is retracted a distance L₈. Next, acollagen hemostasis device 480 or the like is deployed from the centrallumen 59 of electrosurgical catheter 100. When ejected from theconfining lumen 59, collagen hemostasis device 480 expands to fill thefull diameter of the artificial channel 264 over length L₈ as shown inFIG. 22 to effect a seal to prevent blood loss through the opening inthe epicardium 268. Alternatively, suturing techniques may be employed,either percutaneously, thoracoscopically or in an open chest procedureto seal the opening of the artificial channels at the surface of theepicardium 268.

The stent frame 372 of the present invention is typically manufacturedfrom a tubular material, such as tubing made out of shape memory alloyhaving elastic or pseudo-elastic properties, such as Nitinol™, Elgiloy™,or the like. Alternatively, the stent frame may comprise malleablematerials other than shape memory alloys, such as stainless steel. Inthis configuration, the stent frames will preferably be expanded at thetarget site by conventional methods, e.g., an expandable balloon at thedistal end of a catheter shaft. The tubular member is usuallysignificantly smaller in diameter as compared to the final diameter ofthe stent in the expanded configuration within the body lumen. Slots maybe cut into the tubes via laser cutting methods, photo etching, or otherconventional methods to form the separate stent frames. For example,these methods include coating the external surface of a tube withphotoresist material, optically exposing the etch pattern using a laserbeam while translating and rotating the part, and then chemicallyetching the desired slot pattern of the state using conventionaltechniques. A description of this technique can be found in U.S. Pat.No. 5,421,955 to Lau, the complete disclosure which is incorporatedherein by reference. In other methods, laser cutting technology is usedin conjunction with computer controlled stages to directly cut a patternof slots in the wall of the hypodermic tubing to obtain the desiredstent geometry. A description of a typical laser cutting method isdisclosed in U.S. Pat. No. 5,345,057 to Muller, the complete disclosureof which is incorporated herein by reference.

In an exemplary configuration, the stent frame 372 is formed from aresilient shape memory alloy material that is capable of being deformedby an applied stress, and then recovering to its original unstressedshape. The alloy material will usually exhibit thermoelastic behavior sothat the stents will transform to the original unstressed state upon theapplication of heat (i.e., an A_(f) temperature below body temperature).The stents may also exhibit stress-induced martensite, in which themartensite state is unstable and the prosthesis transforms back to theoriginal state when a constraint has been moved (i.e., when the stent isreleased from an introducing catheter within a body lumen). The materialfor the shape memory alloy will be selected according to thecharacteristics desired of the population of prostheses. Preferably, theshape memory alloy will comprise a nickel titanium based alloy (i.e.,Nitinol™), which may include additional elements which affect thecharacteristics of the prosthesis, such as the temperature at which theshape transformation occurs. For example, the alloy may incorporateadditional metallic elements, such as copper, cobalt, vanadium,chromium, iron or the like.

It should be noted that the stents 370 described above and shown inFIGS. 17 and 22 are only representative of the lumenal prostheses thatmay be used with the present invention. The present invention mayincorporate a variety of representative conventional stent structuresmade from metallic tubular materials that are currently marketed asimplants for coronary, peripheral, biliary and other vessels includingthe Palmaz-Schatz™ balloon expandable stent, manufactured by Johnson andJohnson Interventional Systems, Co. and the Memotherm™ stentmanufactured by Angiomed, a division of C.R. Bard, Inc. Other stent orgraft designs that can be incorporated into the present inventioninclude a coiled structure, such as that described in U.S. Pat. No.5,476,505 to Limon, an open mesh or weave stent structure formed ofhelically wound and/or braided strands or filaments of a resilientmaterial, described in U.S. Pat. No. 5,201,757 to Heyn, a filamentknitted into a mesh cylinder, described in U.S. Pat. No. 5,234,457 toAndersen, a tubular structure having diamond shaped openings, describedin U.S. Pat. Nos. 5,242,399 to Lau or U.S. Pat. No. 5,382,261 to Palmaz,Z-shaped stents as described in U.S. Pat. No. 5,282,824 to Gianturco,continuous wire stents, such as the one described in U.S. Pat. No.5,292,331 to Boneau, stents formed of filaments that are wound intospiral or other suitable shapes as described in U.S. Pat. No. 5,314,471to Fountaine, a continuous helix of zig-zag wire and loops described inU.S. Pat. No. 5,405,377 to Cragg and a variety of other types of stents.

1. A method of revascularizing a portion of a patient's myocardiumcomprising: positioning an active electrode in close proximity to atarget site on a wall of the patient's heart wherein the activeelectrode comprises an electrode array including a plurality of isolatedelectrode terminals; contacting the active electrode with anelectrically conducting fluid disposed in a space between the activeelectrode and the target site; inducing discharge of energetic electronsand photons from the conducting fluid by applying a sufficienthigh-frequency voltage between the active electrode and a returnelectrode; and directing the energetic electrons and photons to ablatetissue at the heart wall to form a revascularizing channel through atleast a portion of the heart wall.
 2. The method of claim 1, furthercomprising axially translating the active electrode surface through theportion of the heart wall to form the revascularizing channel.
 3. Themethod of claim 1, further comprising: introducing at least a distal endof an electrosurgical catheter into the ventricle of the heart; andpositioning the distal end of the catheter in close proximity to theendocardium.
 4. The method of claim 1, further comprising: introducingat least a distal end of an electrosurgical probe through an opening inthe patient's chest cavity; and positioning the distal end of the probein close proximity to the epicardium.
 5. The method of claim 4, whereinthe probe is introduced through an intercostal penetration in thepatient.
 6. The method of claim 1, wherein the voltage is appliedcontinuously between the active and return electrodes.
 7. The method ofclaim 1, wherein the voltage is applied in pulses to correspond tobeating of the patient's heart.
 8. The method of claim 1, wherein theactive electrode comprises a single electrode protruding from a distalend of an electrosurgical probe.
 9. The method of claim 1, furtherincluding independently controlling current flow from at least two ofthe electrode terminals based on impedance between the electrodeterminal and the return electrode.
 10. The method of claim 1, furthercomprising forming a revascularizing channel with a lateral dimension ofabout 1.5 to 3.0 mm.
 11. The method of claim 1, further comprisingpositioning a radially expandable lumenal prosthesis in therevascularizing channel to maintain patency of the channel.
 12. Themethod of claim 1, wherein the channel formed by the active electrodesurface is curved.
 13. The method of claim 12, wherein the channelformed by the active electrode surface has first and second openings onone side of the heart wall, and a substantially U-shape therebetween.14. The method of claim 1, wherein the electrode terminals are embeddedin an insulating matrix to electrically isolate each terminal, theinsulating matrix comprising an inorganic material.
 15. The method ofclaim 1, wherein the plurality of return electrodes are proximallyrecessed from the active electrode terminals.
 16. The method of claim 1,wherein the return electrode and the active electrode terminals aredisposed on a distal surface of an electrosurgical probe.
 17. The methodof claim 1, further comprising controlling the depth of therevascularizing channel.
 18. The method of claim 17, further comprisingvisually marking the target site on the heart wall.
 19. The method ofclaim 17, further comprising determining a thickness of the heart wallat the target site.
 20. The method of claim 17, further comprisingsetting a predetermined distance through the heart wall at the targetsite and interrupting the flow of voltage to the active electrodesurface when said active electrode surface has advanced thepredetermined distance to control the depth of the channel.
 21. Themethod of claim 19, wherein the determining step comprises measuringtissue impedance beyond the distal end of the active electrode surface.22. The method of claim 1, further comprising the step of determiningwhen the active electrode surface has substantially penetrated throughthe heart wall.
 23. The method of claim 22, further comprisingterminating the high frequency voltage before the active electrodesurface pierces an opposite wall surface of the heart wall.
 24. Themethod of claim 1, wherein the active electrode is axially translatedthrough at least a portion of the heart wall at a substantially constantrate.
 25. The method of claim 1, wherein the electrically conductivefluid provides a conductive pathway between the active electrode and thereturn electrode.
 26. The method of claim 1, wherein the electricallyconductive fluid is isotonic saline.
 27. The method of claim 1, whereinthe channel is formed by a volumetric removal of the target tissue. 28.The method of claim 1, wherein the applied high-frequency voltagecomprises a peak-to-peak voltage between 40 to 4000 volts.
 29. Themethod of claim 1, wherein the applied high-frequency voltage comprisesa peak-to-peak voltage between 100 to 3200 volts.
 30. The method ofclaim 1, wherein the applied high-frequency voltage comprises apeak-to-peak voltage between 300 to 2400 volts.
 31. A method ofrevascularizing a portion of a patient's myocardium comprising:positioning an active electrode in close proximity to a target site on awall of the patient's heart, wherein the active electrode comprises anelectrode array including a plurality of isolated electrode terminals;directing an electrically conducting fluid in a space between the activeelectrode and the target site; and applying high-frequency voltagebetween the active electrode and a return electrode to ablate tissue atthe heart wall to form a revascularizing channel through at least aportion of the heart wall.